Apparatus and method for estimating stroke volume of the heart using bio-impedance techniques

ABSTRACT

A method of estimating stroke volume of the heart is described. In this method, the volume of the heart is estimated from electrical impedance data of the chest, at two different phases of the cardiac cycle. The stroke volume is estimated from the difference between the volumes estimated at the two phases.

RELATED APPLICATIONS

This application is a divisional of U.S. patent application Ser. No.11/023,912 filed on Dec. 27, 2004, which is a continuation-in-part (CIP)of U.S. patent application Ser. No. 10/189,161 filed on Jul. 3, 2002,now U.S. Pat. No. 7,096,061.

The contents of all of the above applications are incorporated byreference as if fully set forth herein.

FIELD OF THE INVENTION

The present invention relates to the field of instrumentation formonitoring and evaluating patients with heart disease, particularlycongestive heart failure.

BACKGROUND OF THE INVENTION

Congestive heart failure (CHF) is a condition in which the heart doesnot adequately maintain circulation of blood. It is characterized by anincrease in retained body water, especially extracellular water, oftenin the lungs (pulmonary edema). A decrease in extracellular fluid in CHFpatients typically indicates an improvement in heart performance.Conventional methods of monitoring CHF patients either require expensiveequipment and trained personnel (e.g. measuring pulmonary artery andcentral venous pressure with catheters, measuring blood flow through themitral annulus and pulmonary veins with doppler echocardiography) or arenot very accurate (e.g. monitoring changes in body weight, observingneck vein distension, measuring ankle dimensions). Impedancemeasurements of the chest, both resistive and reactive (capacitive)impedance, have been shown to correlate with total body water,extracellular body water, and the ratios of these quantities to fat freemass (U.S. Pat. No. 5,788,643). Monitoring trends in these quantities incongestive heart failure patients is a particularly useful way todetermine whether medication doses need to be increased or decreased. Asstated in U.S. Pat. No. 5,788,643: “Subramanyan, et al. and others haveshown that both the resistive and reactive components of the body'simpedance to the flow of relatively high frequency (50 kHz) electricalcurrent is sensitive to the amount of fluid retained by a patient withCHF. As the CHF resolves, resistance and reactance both increase as doesthe [ratio of reactance to resistance]. See Subramanyan, et al., “TotalBody Water in Congestive Heart Failure,” Jour. Asso. Phys. Ind., Vol.28, September, 1980, pages 257-262 . . . . It would be most desirable toprovide a simple way of detecting increases in body water of patientswith CHF before hospitalization is necessary and permitting adjustmentsin medication and/or diet in time to prevent an episode of acute heartfailure.” The patent describes a figure of merit, calculated fromimpedance measurements, for deciding when medical intervention may beneeded for a CHF patient.

There are several parameters that affect the impedance of the thorax.The impedance of the chest cavity is small compared to changes in theimpedance of the skin, and chest cavity impedance changes substantiallyduring the respiratory cardiac cycle, due to the changing volume of airin the lungs, and during the cardiac cycle due to the changing bloodperfusion of the lungs. Various techniques are used to separate out thepart of the impedance due to excess body water, and to meaningfullycompare such impedance measurements taken in the same patient ondifferent days. For example, U.S. Pat. No. 5,749,369, and Charach, G. etal., “Transthoracic Monitoring of the Impedance of the Right Lung inPatients with Cardiogenic Pulmonary Edema,” Crit. Care Med. 2001, Vol.29, No. 6, pages 1137-1144 discuss ways to compensate for drifting skinimpedance.

In addition to the techniques used in bulk measurements of impedance,impedance imaging is also useful for separating out the differentcontributions to the impedance. In impedance imaging, a set of manyelectrodes (usually 16 or 32) is placed on the body, for exampleencircling the chest, and the voltage is measured at each electrode,while a known current is applied between different pairs of theelectrodes. The resulting data is used to produce a map of the internalimpedance of the body, using various mathematical techniques, some ofthem similar to those used in x-ray tomography. Some imagereconstruction techniques are described in a review paper by D. C.Barber, Med. Phys., (1989), Vol. 16, pages 162-169.

The finite element method, finite difference method, and boundaryelement method are different techniques used to solve differentialequations numerically. Solving Poisson's equation to find the potentialdistribution in the body due to known current sources and impedancedistribution, together with boundary conditions, is known as the forwardproblem. These numerical methods are used in the field of bio-impedanceto solve the forward problem. Rosenfeld, M. et al., “Numerical Solutionof the Potential Due to Dipole Sources in Volume Conductors WithArbitrary Geometry and Conductivity,” IEEE Transactions on BiomedicalEngineering, July 1996, Vol. 43, No. 7, pages 679-689 use a differenttechnique, the finite volume method, to solve the forward problem. Thefinite volume method is also used to solve the forward problem by GuoyaDong et al, “Derivation from current density distribution toconductivities based on the adjoint field theory and numerical test withfinite volume method,” presented at the 2^(nd) Japan, Australia and NewZealand Joint Seminar, 24-25 Jan. 2002, Kanazawa, Japan, on Applicationsof Electromagnetic Phenomena in Electrical and Mechanical Systems.Finding the impedance distribution with known potential distribution atthe surface (measured with surface electrodes, for example), and knowncurrent sources (flowing from one surface electrode to another), iscalled the inverse problem. Some of the inverse problem solvers use theforward problem solver as a step in an iterative solution.

An early paper on impedance imaging by Eyuboglu, B. M. et al., “In VivoImaging of Cardiac Related Impedance Changes,” March 1989, IEEEEngineering in Medicine and Biology Magazine, Vol. 8, pages 39-45discusses the use of gating and time-averaging to separate out thecontributions of the respiratory and cardiac cycles to the chestimpedance and impedance images, including impedance images of pulmonaryembolisms. The authors state, “[T]he resistivity of most tissue changessignificantly with blood perfusion into the tissue . . . . [I]t has beenshown that the thoracic resistivity changes during the cardiac cycle canbe imaged by ECG-gated EIT [electrical impedance tomography] . . . . Theaverage resistivity of lung tissue increases with the amount of airinspired . . . [by] approximately 300 percent . . . from maximalexpiration to maximal inspiration . . . . The resistivity of lung tissuealso changes with the perfusion of blood following ventricular systole .. . . This change has been calculated as 3 percent . . . [which] may beas small as the noise level . . . . Therefore, to pick up thecardiac-related resistivity variations within the thorax during normalbreathing, the respiratory component and the noise must be eliminated .. . . The respiratory component may be rejected by temporal averaging .. . . Experience has shown that averaging over at least 100 cardiaccycles is needed during shallow breathing to attenuate the respiratorycomponent and to improve S/N ratio. Cardiac gating is required . . . .”Brown and Barber develop numerical methods to reduce noise in U.S. Pat.No. 5,311,878, and they use differences in impedance at differentelectrical frequencies between 10 kHz and 600 kHz to distinguish betweencardiac and respiratory effects in U.S. Pat. No. 5,746,214. Newell, J.C. et al., “Assessment of Acute Pulmonary Edema in Dogs by EletricalImpedance Imaging,” February 1996, IEEE Transactions on BiomedicalEngineering, Vol. 43, No. 2, pages 133-138 demonstrate the use ofimpedance imaging to detect pulmonary edemas in dogs, and discuss thevariability in impedance over time and from day to day, which makes itdifficult to measure long-term changes.

The disclosures of the patents and the papers listed above areincorporated herein by reference.

SUMMARY OF THE INVENTION

An aspect of some embodiments of the invention concerns the use of anelectrocardiograph (ECG) to measure the depth, frequency, and/or timingof the breathing cycle, in order to be able to correct for the effect ofbreathing on the chest impedance, which would otherwise mask the effectsof pulmonary edema and other symptoms of congestive heart failure on thechest impedance. The breathing cycle is correlated with the RR Intervalsextracted from ECG data, because breathing modulates the heart'spacemaker located at the sinuatrial node. Breathing depth also affectsthe amplitude of the raw ECG data, since the higher impedance of thechest when the lungs are expanded reduces the voltage at the ECGelectrodes. By tracking changes in the ECG data at a given point in thecardiac cycle, for example the minimum voltage or the maximum voltagebetween electrodes during each cardiac cycle, the breathing cycle can bemonitored. Although the breathing cycle can also be monitored directly,by measuring air flow into and out of the lungs, this requires morepatient cooperation than taking ECG data does, and requires extraequipment, so it is easier to monitor breathing by using ECG data. ECGdata is usually obtained anyway in impedance imaging, in order tomonitor the cardiac cycle, and no extra equipment is needed if the ECGdata is used to monitor the breathing cycle at the same time.Optionally, the system is adapted to be used as home monitoring system,with the information transferred to a remote location where a physicianviews and diagnoses the condition of a patient. The data can betransferred, for example, by a modem over telephone lines, throughsecure broadband internet lines, or by another means of communication.

An aspect of some embodiments of the invention concerns solving theinverse problem, i.e. calculating an impedance image of the chest frommeasured voltages between different pairs from a set of electrodes onthe surface of the body, using the finite volume method. The finitevolume method offers several advantages over the finite element methodand boundary element method for solving the inverse problem, but it hasnot previously been used for solving the inverse problem in impedanceimaging.

An aspect of some embodiments of the invention concerns using ECG data,together with impedance imaging, to evaluate the condition of acongestive heart failure patient, for example in order to determinewhether to increase or decrease doses of medication. Diuretics, forexample, which are prescribed to reduce pulmonary edema and othersymptoms of congestive heart failure, may induce cardiac arrhythmia iftaken in too high a dose. In determining the optimal dose, patientoutcome is likely to be better if treatment is determined by looking atthe overall picture, including symptoms of congestive heart failure andsymptoms that may indicate incipient arrhythmia, as well as othersymptoms that may be seen in ECG data, rather than simply starting orstopping medication based on isolated symptoms. U.S. Pat. No. 5,788,643describes a figure of merit for deciding when medical intervention iscalled for in a CHF patient, but this figure of merit is based only onimpedance measurements, not on ECG data.

Optionally, the ECG data is also used to measure the breathing cycle tocorrect the impedance imaging, as described above. Optionally, theelectrodes used for the ECG are also used for the impedance imaging.

An aspect of some embodiments of the invention concerns using impedanceimaging to measure the stroke volume of the heart. Impedancemeasurements are optionally made at the time of end-systole andend-diastole, as determined, for example, by an ECG. The impedancemeasurements are used to make a best fit to the dimensions of theinterior of the heart, which has a high conductivity, and hence toestimate the volume of the interior of the heart, at the two phases. Thedifference in volumes is a good approximation to the stroke volume.

Optionally, the impedance measurements, whether they are used to measurefluid in the lungs, or to measure stroke volume of the heart, or foranother purpose, are performed using an automatic system suitable foruse in hospitals.

There is thus provided, in accordance with an embodiment of theinvention, a method for generating impedance images of the chest,comprising:

acquiring electrical data of the chest;

obtaining electrocardiograph data of a patient;

analyzing the electrocardiograph data to obtain information aboutbreathing parameters at the time the electrical data was acquired; and

reconstructing at least one impedance image of the chest from theelectrical data and the information about breathing parameters;

wherein the information about breathing parameters reduces thesensitivity of the at least one impedance image to breathing parameters.

Optionally, reconstructing at least one impedance image comprises:

reconstructing at least one preliminary impedance image of the chestfrom the electrical data; and

correcting the at least one preliminary impedance images to form the atleast one impedance image, taking into account the breathing parameters.

Optionally, analyzing the electrocardiograph data comprises analyzingchanges in RR intervals.

Alternatively or additonally, analyzing the electrocardiograph datacomprises analyzing changes in a voltage measured at a same phase ineach cardiac cycle.

Alternatively or additionally, analyzing the electrocardiograph datacomprises analyzing the average over one or more cardiac cycles of avoltage measured by the electrocardiograph.

In an embodiment of the invention, reconstructing at least onepreliminary image comprises reconstructing a plurality of preliminaryimages, and correcting the at least one impedance images comprisessorting the preliminary images into a plurality of bins according to thebreathing parameters.

Optionally, sorting the preliminary images into bins comprises sortingaccording to the state of expansion of the lungs.

Alternatively or additionally, sorting the preliminary images into binscomprises sorting according to the elapsed time since the last maximumexpansion of the lungs.

Alternatively or additionally, sorting the preliminary images into binscomprises sorting according to the elapsed time since the last minimumexpansion of the lungs.

Optionally, sorting the preliminary images into bins comprises sortingaccording to a cardiac volume.

Alternatively or additionally, sorting the preliminary images into binscomprises sorting according to a heart rate.

Alternatively or additionally, sorting the preliminary images into binscomprises sorting according to a phase of the cardiac cycle.

In an embodiment of the invention, acquiring the electrical datacomprises gating by the cardiac cycle.

Optionally, gating by the cardiac cycle comprises using the peak of theR-wave to trigger acquiring the electrical data.

Optionally, correcting the at least one preliminary impedance imagescomprises averaging the impedance data acquired over one or morebreathing cycles.

Alternatively or additionally, reconstructing at least one preliminaryimage comprises reconstructing a plurality of preliminary images forwhich the impedance data was acquired at a plurality of phases in thebreathing cycle, and correcting the at least one preliminary impedanceimages comprises averaging the preliminary impedance images.

Optionally, the method includes measuring the air flow into the lungs,and calibrating the information about breathing parameters obtained fromthe electrocardiograph using said measured air flow.

Alternatively or additionally, the method includes measuring the airflow out of the lungs, and calibrating the information about breathingparameters obtained from the electrocardiograph using said measured airflow.

Optionally, reconstructing at least one preliminary impedance image ofthe chest comprises using a finite volume method.

There is further provided, according to an embodiment of the invention,a method for generating an impedance image of the chest, comprising:

acquiring electrical data of the chest; and

using a finite volume method to calculate an impedance image from theelectrical data.

Optionally, the method includes:

formulating an initial impedance image;

using a finite volume method to calculate an expected set of electricaldata if the impedance distribution of the chest matched the initialimpedance image;

determining a difference between the acquired electrical data and theexpected electrical data; and

calculating a new impedance image based on said difference.

Optionally, calculating an expected set of electrical data andcalculating a new impedance image are iterated at least one time, usingthe new impedance image calculated in at least one previous iteration tocalculate the expected set of electrical data in each iteration exceptthe first iteration.

Optionally, calculating an expected set of electrical data andcalculating a new impedance image are iterated until the differencebetween the acquired electrical data and the expected set of electricaldata is small enough to satisfy a stopping condition.

Optionally, calculating the new impedance image comprises calculatingwith a Newton-Raphson method.

Alternatively or additionally, calculating the new impedance imagecomprises calculating with a modified Newton-Raphson method.

In an embodiment of the invention, formulating the initial impedanceimage comprises ascribing typical impedances to different parts of thechest according to at least one image of the chest.

Optionally, ascribing impedances according to at least one image of thechest comprises ascribing impedances according to at least one x-rayimage.

Optionally, ascribing impedances according to at least one x-ray imagecomprises ascribing impedances according to at least one x-ray computedtomography image.

Alternatively or additionally, ascribing impedance according to at leastone image of the chest comprises ascribing impedances according to atleast one magnetic resonance image.

Alternatively or additionally, ascribing impedances according to atleast one image of the chest comprises ascribing impedances according toat least one ultrasound image.

In an embodiment of the invention, using the finite volume methodcomprises inverting a matrix with a technique that is adapted forinverting sparse matrixes.

Optionally, inverting a matrix comprises inverting a matrix with thesuccessive over relaxation method.

Optionally, acquiring electrical data of the chest comprises measuringpotentials at a plurality of locations on the body, while known currentsare applied at a plurality of locations on the body.

Optionally, applying known currents comprises applying a current betweena first pair of current-applying locations at substantially oppositesides of the chest.

Optionally, measuring potentials comprises measuring the potentialdifference between two voltage-measuring locations at substantiallyopposite sides of the chest, different from the first pair ofcurrent-applying locations.

Optionally, applying known currents is repeated, using a second pair ofcurrent-applying locations that are substantially on opposite sides ofthe chest and differ from the first pair of current-applying locations.

Optionally, one of the pairs of current-applying locations comprise alocation on the left side of the front of the chest and a location onthe right side of the back, and the other pair of current-applyinglocations comprise a location on the right side of the front of thechest and a location on the left side of the back.

There is further provided, in accordance with an embodiment of theinvention, a method of estimating stroke volume of the heart,comprising:

generating a first impedance image of the chest according to anembodiment of the invention, at a first phase of the cardiac cycle;

estimating a first volume of the heart from the first impedance image;

generating a second impedance image of the chest according to anembodiment of the invention, at a second phase of the cardiac cycle;

estimating a second volume of the heart from the second impedance image;and

using the difference between the first and second volumes of the heartto estimate the stroke volume of the heart.

There is further provided, in accordance with an embodiment of theinvention, a method of estimating stroke volume of the heart,comprising:

taking a first set of electrical data of the chest at a first phase ofthe cardiac cycle;

estimating a first volume of the heart from the first set of electricaldata;

taking a second set of electrical data of the chest at a second phase ofthe cardiac cycle;

estimating a second volume of the heart from the second set ofelectrical data; and

using the difference between the first and second volumes of the heartto estimate the stroke volume of the heart.

Optionally, estimating at least one of the volumes of the heart from thecorresponding set of electrical data comprises generating an impedanceimage of the chest from said set of electrical data.

Optionally, one of the first and second phases of the cardiac cyclecomprises an end-systole phase.

Optionally, one of the first and second phases of the cardiac cyclecomprises an end-diastole phase.

Optionally, for each of said phases of the cardiac cycle, generating theimpedance image comprises generating a two-dimensional impedance imageof a slice of the chest, and estimating the volume of the heartcomprises estimating a cross-sectional area of the interior of theheart.

Optionally, generating the impedance image comprises modeling thecross-sectional area of the interior of the heart as an ellipse.

Optionally, estimating the volume of the heart comprises modeling thecross-sectional area of the interior of the heart as an ellipse.

Optionally, estimating the volume of the heart comprising using aformula which gives the volume of the heart as a function of thecross-sectional area.

Optionally, the volume is proportional to the square of thecross-sectional area.

Optionally, estimating the first volume of the heart comprisesgenerating a first impedance image of the chest, and estimating thesecond volume of the heart comprises generating a second impedance imageof the chest putting constraints on differences between the secondimpedance image and the first impedance image, but not constraining thevolume of the heart to be the same in the first and second images.

Optionally, the second impedance image, aside from the heart, isconstrained to be the same as the first impedance image.

In an embodiment of the invention, acquiring electrical data of thechest comprises measuring potential differences between one or morepairs among a plurality of voltage-measuring locations on the body,while applying known currents at a plurality of current-applyinglocations on the body.

Optionally, each of the pairs of current-applying locations comprises alocation on the left side of the front of the chest, and a location onthe right side of the back.

Optionally, for each time the known currents are applied, the pluralityof voltage-measuring locations comprises three different locations, atleast two of them on the front of the chest, and all three of themdifferent from either of the pair of current-carrying locations at whichthe known currents are being applied at that time.

There is further provided, in accordance with an embodiment of theinvention, a method for monitoring a congestive heart failure patient,comprising:

generating at least one impedance image of the patient's chest;

acquiring electrocardiograph data of the patient; and

calculating a parameter characterizing medical treatment of the patient,from electrocardiograph data and at least one impedance image of thechest.

Optionally, calculating at least one parameter comprises calculating arecommended dose of a medication.

Optionally, calculating a recommended dose of medication comprisescalculating a recommended dose of a diuretic.

Optionally, using the electrocardiograph data comprises using the QTinterval.

Optionally, using the QT interval comprises using the QT interval todetect hypokalemia.

Alternatively or additionally, using the electrocardiograph datacomprises using the U wave amplitude.

Optionally, using the U wave amplitude comprises using the U waveamplitude to detect hypokalemia.

There is further provided, in accordance with an embodiment of theinvention, an apparatus for making corrected impedance images of thechest, comprising:

an impedance imaging data acquisition system which acquires impedanceimaging data of the chest;

an electrocardiograph which obtains electrocardiograph data of apatient; and a data analyzer which analyzes the electrocardiograph datato obtain information about breathing parameters at the time theimpedance imaging data was acquired, and reconstructs, from theimpedance imaging data and the information about breathing parameters,at least one impedance image of the chest with reduced sensitivity tobreathing parameters.

There is further provided, in accordance with an embodiment of theinvention, an apparatus for making impedance images of the chest,comprising:

an impedance imaging data acquisition system which acquires impedanceimaging data of the chest; and

a data analyzer which reconstructs an impedance image of the chest fromsaid impedance imaging data, using a finite volume method.

There is further provided, in accordance with an embodiment of theinvention, an apparatus for estimating the stroke volume of the heart,comprising:

an electrocardiograph;

an impedance data acquisition system which acquires impedance data ofthe chest;

a controller which uses data from the electrocardiograph to trigger thedata acquisition system to acquire the data at each of a first phase anda second phase of the cardiac cycle; and

a data analyzer which estimates the volume of the heart at the firstphase and the second phase from the data acquired at the first phase andthe second phase, thereby allowing the stroke volume to be estimatedfrom the difference between the volume of the heart at the first phaseand the second phase.

Optionally, the data analyzer is configured to analyze theelectrocardiograph data to obtain information about breathing parametersat the time the impedance imaging data was acquired, and is configuredto use the information about breathing parameters in estimating thevolume of the heart at the first phase and the second phase, therebymaking the estimate of the stroke volume more accurate.

Optionally, the electrocardiograph, impedance data acquisition system,and controller comprise a self-contained portable system weighing lessthan 5 kilograms.

Optionally, there is also a user interface to the controller whereby theuser initiates a pre-set sequence of acquisition of the impedance data,the time of the acquisition being triggered by data from theelectrocardiograph.

BRIEF DESCRIPTION OF THE DRAWINGS

Exemplary embodiments of the invention are described in the followingsections with respect to the drawings. The drawings are generally not toscale. Features found in one embodiment can also be used in otherembodiments, even though not all features are shown in all drawings.

FIG. 1 is schematic view of a cross-section of the chest, showing theplacement of electrodes for impedance imaging, according to prior art;

FIG. 2 is a flowchart showing how ECG data is used to distinguish theeffect of breathing from the effect of the cardiac cycle on an impedanceimage of the chest, according to an exemplary embodiment of theinvention;

FIGS. 3A, 3B and 3C show breathing data and ECG data, illustrating howthe ECG data is affected by breathing;

FIG. 4 is a schematic drawing of a hardware configuration for impedanceimaging, according to an exemplary embodiment of the invention;

FIG. 5 is a flowchart showing how the finite volume method is used tocalculate an impedance image, according to an exemplary embodiment ofthe invention;

FIG. 6 is a flowchart showing how ECG data and impedance images are usedto assess the condition of a congestive heart failure patient, accordingto an exemplary embodiment of the invention;

FIGS. 7A and 7B are schematic cross-sectional views of the chest, atend-systole and end-diastole phases respectively, according to anexemplary embodiment of the invention;

FIG. 8 is a perspective view of a hospital bio-impedance system,according to an exemplary embodiment of the invention;

FIG. 9 is a schematic cross-sectional view of the chest, showing theplacement of 8 electrodes, using the bio-impedance system shown in FIG.8; and

FIG. 10 is a flow chart showing a procedure for using the bio-impedancesystem shown in FIG. 8.

DETAILED DESCRIPTION OF EXEMPLARY EMBODIMENTS

Aspects of some embodiments of the invention concern systems for makingimpedance images of the chest, and for using these images to monitorcongestive heart failure patients, or to estimate the stroke volume ofthe heart. In order to describe the embodiments of the invention shownin FIGS. 2-10, it will be convenient to first describe some prior artshown in FIG. 1. The various options described for FIG. 1 are alsooptions for the embodiments of the invention shown in FIGS. 2-10.

FIG. 1 shows a cross-section of a chest 100, including lungs 102 and aheart 104. An impedance data acquisition system comprises sixteenelectrodes 106, shown placed on the skin all around the chest. Althoughnot shown in FIG. 1, the impedance data acquisition system also includesa current source such as a power supply, for passing current through theelectrodes, and optionally an amplifier for amplifying voltagemeasurements made by the electrodes. The number of electrodes used isoptionally great enough to obtain a desired resolution in the impedanceimage, but not so great that the measurements and data analysis take toolong. Eight, sixteen and thirty-two are numbers that are commonly used,but other numbers of electrodes may be used. Powers of two have thepotential advantage that they are generally more efficient to use forfast fourier transform (FFT) algorithms. To take a set of electricaldata for an impedance image, current is first passed through two of theelectrodes, and the voltage is measured at all of the electrodes. Thenanother pair of electrodes is chosen for passing current through, andthe process is repeated for many different pairs of electrodes.Optionally, the voltage is not measured on the electrodes with currentpassing through them, since for those electrodes the voltage tends to bedominated by the voltage drop between the electrode and the skin, so itis difficult to obtain accurate potential measurements on thoseelectrodes. Optionally, more than one pair of electrodes has currentpassing through it, for one or more of the measurements. In this case,different electrodes optionally have different currents flowing throughthem. Although this may make the data analysis simpler, it has thedisadvantage that there are more electrodes for which it is difficult toget good potential measurements. Optionally, one or more of theelectrodes are also used to obtain ECG data.

In FIG. 1, the electrodes are arranged in a single circle around thebody, similar to the arrangement used by Eyuboglu, Brown and Barber(loc. cit.). This arrangement may not provide any information about theaxial distribution of impedance inside the body, but provides atwo-dimensional cross-sectional map of impedance, a weighted averageover the axial direction of the three-dimensional impedancedistribution. Optionally, the electrodes are arranged not in a singlecircle, but in two or more circles at different axial positions. Such atwo-dimensional grid of electrodes provides data for constructing athree-dimensional map of impedance. More than one circle of electrodesis optionally used for other reasons as well. For example, optionallythe positive electrode supplying current is always located in onecircle, and the negative electrode with current is always in the othercircle. This arrangement provides more independent measurements than ifthe positive and negative electrodes were chosen from the same circle ofelectrodes, since in that case switching the two electrodes would notprovide any new information. Having one circle of electrodes forpotential measurements, and one or two separate circles of electrodesfor supplying current, also avoids the problem of measuring potential onan electrode that is supplying current.

Typical currents used for impedance imaging are 1 to 5 milliamps. Acurrent of this magnitude is not dangerous, but is high enough toprovide a reasonable signal to noise ratio when measuring the voltage.In order to obtain reactive (capacitive) impedance data as well asresistance data, the currents optionally are AC, typically atfrequencies between 10 kHz and several hundred kHz. However, lowerfrequencies may also be used. For safety reasons, DC current istypically not used in medical procedures, even if reactive impedancedata is not needed. Reactive impedance is related to the capacitance ofcell membranes, and resistive impedance is related to the volume ofwater. Because low frequency currents cannot penetrate the cellmembranes, low frequency resistive impedance tends to measure only thevolume of extracellular water, while high frequency resistive impedancemeasures the volume of water within cells as well.

FIG. 2 is a flowchart describing a procedure for using ECG data tomonitor the state of expansion of the lungs, and to calibrate impedanceimages of the chest according to the state of expansion of the lungs.Using this procedure, it may be possible to detect the relatively smallchanges in impedance associated with changes in thoracic fluid volume,in spite of the larger changes in impedance associated with breathing.

At 202, a pair of electrodes is chosen to apply current. At 204, thevoltage is measured and recorded on each electrode, while current isflowing through the chosen electrodes. Optionally, as discussed above,the voltage is not measured on the electrodes carrying current, orcertain electrodes are dedicated to carrying current and otherelectrodes are dedicated to measuring the potential. At 206, the flowgoes back to 202 and another pair of electrodes is chosen to carrycurrent, until data has been taken with every possible pair ofelectrodes, or until it is decided, based on some criterion, that asufficient set of data has been taken. The potential data is thenstored, at 208, together with ECG data taken at the same time. At 210,the procedure goes to 212, and a new set of potential measurements isinitiated, until it is decided that a sufficient number of data setshave been taken. Optionally, data sets are taken at intervals shortcompared to the cardiac cycle time, and data is taken over a periodcorresponding to several breathing cycles, at least. This allows theimpedance images to be correlated with the cardiac and breathing cycles.At 214, after all the data has been taken, an impedance image iscomputed for each data set, and associated with the ECG data taken atthe same time. Optionally, the image is computed using the finite volumemethod, according to the procedure detailed below in the description ofFIG. 4.

At 216, the impedance images are sorted by the phase of the cardiaccycle, and by the state of expansion of the lungs, as indicated by theECG data taken at the same time the impedance data was measured for thatimage. The state of expansion of the lungs is optionally inferred fromone or both of two different features of the ECG data. When the lungsare in a more expanded state, the RR interval increases, since theexpansion of the lungs affects the heart's pacemaker located at thesinuatrial node. Optionally, in using the RR interval to infer the stateof expansion of the lungs, variations in the RR interval at frequenciesmuch lower than the breathing frequency are filtered out, since thesecould be due to other factors which affect the RR interval, for examplestress. In addition, the expansion of the lungs increases the resistiveimpedance of the chest, and this reduces the voltage measured by the ECGelectrodes. Normally, in ECG systems, the raw voltage signals areadjusted by pre-amps, which compensate for the slow changes in voltageassociated with the breathing cycle, which are not usually of interest.In order to use this aspect of the ECG data to monitor breathing, thepre-amps may be bypassed.

Optionally, the state of expansion of the lungs as inferred from ECGdata is calibrated by direct measurements of lung expansion, for exampleby measuring the air flow into and/or out of the lungs. Optionally, theimpedance images are also sorted into bins by the rate of expansion orcontraction of the lungs, or other characteristics of the breathing thatmay affect the impedance image, especially the appearance of pulmonaryedemas in the impedance image. If the heartbeat is irregular in strengthor timing, then the images are also optionally sorted by systolicvolume, interval of ventricular contraction, and other characteristicsof the heartbeat that may affect the impedance image.

At 218, the sorted impedance images are converted to a canonicalimpedance image in which the appearance of pulmonary edema, or themeasured thoracic fluid volume, is independent of the cardiac andbreathing cycles. At 220, the canonical image is stored. Such acanonical image may be used to meaningfully compare thoracic fluidvolume, or other characteristics of a pulmonary edema, at differenttimes, hours or days or weeks apart, and to detect trends which mayindicate the need to increase or decrease doses of medication, or tostop or start a given medication, or to intervene medically in otherways.

Optionally, instead of computing preliminary impedance images at 214 andthen sorting them at 216, the data sets are sorted at 216, with orwithout some preliminary processing, and the sorted data sets are usedto produce a canonical impedance image at 218. Since the data setscontain the information used to produce the preliminary images, itshould be understood that any manipulations performed on the preliminaryimages to produce a corrected image might instead be performed directlyon the data sets without first producing preliminary images.

Several different concepts may optionally be used, singly or in anycombination, in processing the images to produce a canonical image:

-   -   1. Averaging the images in a given bin (for example, the images        taken at a given state of expansion of the lungs, and a given        phase of the cardiac cycle), and then taking a linear        combination of images in different bins.    -   2. The coefficients of this linear combination may be negative.        For example, if the change in impedance of the lungs associated        with a pulmonary edema is correlated with the cardiac cycle,        then images taken at one phase in the cardiac cycle may be        subtracted from images taken 180 degrees apart in the cardiac        cycle. Such a procedure may emphasize pulmonary edemas in the        resulting canonical image, and de-emphasize other features of        chest impedance that are not of interest.    -   3. Changes in chest impedance at the breathing frequency, which        are likely not to be of interest, are eliminated or reduced by        averaging over bins that represent different phases in the        breathing cycle, at the same phase in the cardiac cycle.    -   4. Converting an image taken at any state of lung expansion to        an equivalent image at a canonical state of lung expansion, for        example with the lungs fully expanded, or the lungs emptied, or        half-way in between. An algorithm which does this could make use        of a series of impedance images taken at different states of        expansion of the lungs.

Optionally, the algorithm for producing a canonical impedance image isadjusted for the particular patient, based on previous data taken forthat patient. Additionally or alternatively, the algorithm is based onprevious data taken from one or more other patients, possibly from alarge number of other patients.

FIG. 3A shows lung volume as a function of time for six breathingcycles, FIG. 3B shows the raw ECG data, and FIG. 3C shows RR intervalderived from the ECG data, plotted for the same time period. When thelungs are more expanded, the chest impedance is greater, and the voltageat the ECG electrodes is lower. Hence there is a negative correlationbetween ECG voltage and lung volume. The RR interval is also correlatednegatively with lung volume, because respiration affects the pacemakerof the heart in the sinuatrial node. The correlations between lungvolume, raw ECG voltage, and RR interval are strong enough so that ECGvoltage and RR interval may be usefully used to monitor the state ofexpansion of the lungs during breathing.

FIG. 4 schematically shows a hardware configuration for an impedanceimaging system which uses ECG data to determine breathing parameters, inaccordance with an embodiment of the invention. The hardware comprises acurrent injection module 609, a potential measuring and processingmodule 611, and a user interface module 625. In the current injectionmodule, a 32.768 kHz oscillator 602 generates a stable sinusoidalcurrent of a few micro-amperes, which is amplified to the desiredcurrent, 1 to 5 milliamperes, by current amplifier 604. A dual 1-to-4multiplexer 606 is used to inject the current through any desired pairchosen from 8 electrodes 608, which are placed around the thorax of ahuman body 610, or around a phantom. Potential measuring and processingmodule 611 includes eight electrodes 612, which are applied to thethorax and sense voltage, analog amplifiers 614, and a Motorola DSP56807chip 616. An electrocardiogram 618 also feeds voltage measurements intochip 616. Chip 616 includes an analog to digital convertor 620 whichconverts the analog voltage data to digital form, a central processingunit 622, and a memory 624. The digital data is stored in the memory,for each pair of electrodes used to inject current, and is then used bythe CPU to reconstruct an impedance image. The CPU also uses the datafrom the ECG to calculate parameters such as RR and QT intervals, whichare used to infer breathing parameters. User interface module 625includes a keypad 626 used to enter data or feedback from the user intothe CPU, a liquid crystal display 628 for presenting the results or forgiving instructions to the patient during the measurement process, and adigital to analog convertor 630 for plotting data during development ofthe system. A 9 volt battery 632 provides power for all three modules,via a battery interface 634, which provides positive and negativevoltage and a ground.

Optionally, user interface module 625 is located remotely, with the datatransmitted (for example, over phone lines with a modem, or over asecure broadband internet connection), or user interface module 625includes hardware for transmitting the impedance imaging data frommemory 624 to a remote location. Optionally, current amplifier 604 andmultiplexer 606 are also controlled remotely, or they are controlled bya computer, optionally chip 616, which is programmed to inject a givensequence of currents through the different electrodes. These options maybe useful, for example, for monitoring the condition of a patient who isat home, without the need for him to come into a hospital every time.

FIG. 5 is a flowchart outlining how the finite volume method is used tocalculate an impedance image from the potential data taken withdifferent pairs of electrodes carrying current. Initially, in 402, animage is made of the chest of the patient, using, for example, magneticresonance imaging, computerized x-ray tomography, or ultrasound.Alternatively, with some loss of accuracy, the patient's chest ismodeled by some standard body model, perhaps parameterized bycharacteristics such as weight, height, gender, and body type.Optionally, the model or image includes the whole body, or more of thebody, rather than just the chest, which makes it possible to moreaccurately account for current paths that are not confined to the chest.

At 404, the chest or body model is used to create a three-dimensionalgrid. Optionally, the grid conforms to the surface of the body.Optionally, the grid conforms to the surfaces of the lungs and/or theheart, which generally have substantially different impedance from otherparts of the chest, and from each other. Optionally, the grid changesduring the breathing cycle and heart beat, so that it can continue toconform to the surfaces of the lungs and heart. Alternatively, the gridconforms only to some approximate average surfaces of the lungs andheart, or does not conform to the surfaces of the lungs and heart atall. The grid coordinates of the various electrodes (including theirorientations and outlines, as well as their positions) are determinedand stored.

In 406, potential data is read at each electrode, for each pair ofcurrent-carrying electrodes, as described above in the description ofFIG. 1 and FIG. 2. In 408, an initial guess is made of the impedancedistribution of the chest, for example, using information about thelocation of the lungs and heart obtained from the image made in 402,and/or from a chest model used in 402. Optionally, the initial guess forthe impedance distribution simply assigns typical values of impedancefor lung tissue, cardiac tissue, and the rest of the chest cavity.

In 410, the finite volume method is used to solve the forward problem,calculating the expected surface potential at each electrode wherevoltage is measured, for each choice of current carrying electrodes,using the initial guess for impedance distribution as a starting point.The finite volume method uses the integral form of Poisson's equation,which becomes a set of simultaneous linear equations when Poisson'sequation is discretized and the integral is replaced by a sum. Theboundary conditions for Poisson's equation are Neumann-type conditions,stating the current flux normal to the boundary.

The finite volume method is more accurate than the finite elementmethod, the most commonly used method in the field of bio-impedance, atsolving Poisson's equation with Neumann boundary conditions, because itcan treat discontinuous impedance distributions and discontinuouscurrent sources (B. Lucquin and O. Pironneau, Introduction to ScientificComputing, John Wiley & Sons, 1998, pp. 300-304), and anisotropicconductivities. Discontinuous impedance distributions are a commonfeature of the human body, with different body tissues found inwell-defined organs, and some body tissues are best modeled byanisotropic conductivities. The finite volume method also makes moreefficient use of computational resources and CPU time than the finiteelement method (Abboud, S. et al, Comput. Biomed. Res., (1994), Vol. 27,pages 441-455). The set of linear equations can be represented in sparsematrix form, and relaxation methods can be used that are very fast andefficient for sparse matrixes, for example the successive overrelaxation (SOR) method. The finite volume method also poses less severerestrictions on the quality of the mesh than the finite element method(H. K. Versteeg and W. Malalasekera, An Introduction to ComputationalFluid Dynamics—The Finite Volume Method, Longman Scientific & Technical,1995). The references cited are incorporated herein by reference. Inspite of these advantages of the finite volume method, the finiteelement method has often been used because of the ready availability ofcommercial forward solvers using the finite element method. The greaterspeed and more efficient use of computational resources by the finitevolume method may be more important when solving the inverse problemthan they are when solving the forward problem.

In 412, the surface potential calculated at each electrode in 410, foreach chosen pair of current-carrying electrodes, is compared to thevoltages measured at each electrode in 406. If the difference betweenthe measured and calculated potentials is small enough, then the initialguess made in 408 for the impedance distribution is a good match to theactual impedance distribution. Otherwise, the Newton-Raphson method or asimilar method may be used in 414 to make an improved guess for theimpedance distribution, and step 410 (solving the forward problem) isrepeated, using the new guess. The Newton-Raphson method involvesdifferentiating (finding the Jacobian of) the matrix associated with theset of linear equations in 410, with respect to changes in the impedancedistribution. Here the finite volume method offers another advantageover the finite element method, since the finite volume method allowsthe matrix elements to be expressed symbolically in terms of theimpedance distribution, and the expressions can be mathematicallymanipulated to find their derivatives, and hence the Jacobian. With thefinite element method, on the other hand, the matrix is found only innumerical form, and finding the Jacobian is then much more timeconsuming, for a large matrix.

The Newton-Raphson method involves inverting a matrix, called theHessian matrix, which depends on the Jacobian and on the differencebetween the measured and calculated potentials. Because the Hessianmatrix is often ill-conditioned, the Newton-Raphson method may beunstable. Optionally, the stability of the convergence is improved byusing a modified Newton-Raphson method, for example the Marquardtmethod. These methods involve adding to the Hessian matrix aregularization matrix, which makes it better conditioned.

At each iteration of the loop shown in FIG. 5, the calculated potentialis compared to the measured voltages on the electrodes. When thedifference between them is small enough, the latest guess for theimpedance distribution is accepted as a good approximation to the actualimpedance distribution. In 416, this impedance distribution is stored,and optionally displayed on a monitor or printed.

FIG. 6 is a flowchart showing how impedance imaging is combined with ECGdata to produce an overall evaluation of a patient suffering fromcongestive heart failure, and to decide on appropriate treatment. ECGdata is recorded in 502. This data is used both for determiningbreathing parameters in 504, as described above in FIG. 2, and fordetecting problems with heart function, for example arrhythmia orincipient arrhythmia, in 506. At the same time, in 508, impedanceimaging is used to estimate the thoracic fluid volume in 510, and thisestimate is adjusted by taking into account the breathing parametersdetermined in 504. This leads in 512 to a canonical impedance image, asdiscussed above in FIG. 2, which characterizes the thoracic fluidvolume, and the presence of pulmonary edema, independently of the stateof expansion of the lungs and the phase of the cardiac cycle at the timethe image was made.

In 514, the canonical impedance image in 512 is used, together with theinformation on cardiac performance in 506, as input to an algorithmwhich generates an evaluation of the patient's overall condition, with aview toward determining the optimal treatment in 516. For example, anabnormally high thoracic fluid volume by itself might indicate the needfor the patient to take an increased dose of diuretic medication. Butsome diuretics, such as thiazide, furosemide, and ethacrynic acid, cancause or enhance hypokalemia, which if not treated can lead toarrhythmia. If the ECG data in 506 shows abnormally long QT intervals,especially with prominent U waves, then this by itself might indicatehypokalemia and the need to decrease the dose of diuretics. Only bylooking at both ECG data in 506 and impedance imaging in 512, is itpossible to determine the optimum dose of medication. An algorithm whichuses both ECG data and impedance imaging, and finds the optimumtreatment, is optionally based, for example, on experience with theoutcomes of other patients with similar combinations of symptoms.

FIGS. 7A and 7B show impedance images of the chest, used to estimate thestroke volume of the heart. The images show a cross-section of thechest, and the impedance data is optionally obtained from electrodes ina belt surrounding the chest, all located approximately in the plane ofthe cross-section. As discussed above in connection with FIG. 1,impedance data from such electrodes can be used to reconstruct animpedance image in a cross-section of the chest corresponding to theplane where the electrodes are located. Alternatively, the electrodesare distributed at different positions longitudinally, as well as aroundthe chest, and are used to construct a three-dimensional impedanceimage. Either a two-dimensional or a three-dimensional impedance imagemay be used to estimate the stroke volume.

FIG. 7A is an impedance image 700 showing the impedance distribution inthe cross-section of the chest, at the end-systole phase of the cardiaccycle, when the left ventricle is fully contracted. The impedance imageis calculated, for example, using any of the methods described above inthe description of the flow chart in FIG. 5, with the impedancemeasurements gated to the end-systole phase of the cardiac cycle, asindicated, for example, by an ECG. The end-systole phase corresponds tothe peak of the R-wave. Optionally, the timing of the end-systole phaseis estimated by recording the ECG signal for at least a few cardiaccycles before taking the impedance data, and measuring the RR interval.The impedance data is then taken at the expected time of the nextend-systole phase.

The impedance image is optionally made using a low enough frequency, forexample 20 kHz, so that it is sensitive mostly to extracellular fluid.(At higher frequencies, for example above 100 kHz, the capacitiveimpedance of the cell membranes is lower relative to the impedance ofthe extracellular and intracellular fluid, so much of the current flowsthrough the cells instead of around them. At such high frequencies, theimpedance of body tissue is less sensitive to the amount ofextracellular fluid, but depends more on the total fluid content.) Thelungs 702 and 704, and the interior 706 of the heart, have particularlylow impedance.

The interior of the heart, in the plane of the image, is optionallymodeled as an ellipse with a principal axis 708 oriented in theanterior-posterior direction, and a principal axis 710 oriented in theleft-right direction. The ellipse has diameter “a” in the direction ofaxis 708, and diameter “b” in the direction of axis 710. Optionally, thecenter of the ellipse is assumed to be at a particular location which isknown to be at least a good approximation to the location of the centerof the heart. Alternatively, a best fit to the location of the center ofthe ellipse is found from the impedance data. This elliptical model is afairly good approximation to the actual cross-sectional shape of theheart in the plane of electrodes, and using such a model has thepotential advantage that a relatively accurate measurement of the heartvolume can be made even with only eight electrodes. Alternativelyanother model is used for the cross-section of the heart, for example itis assumed to be circular, or an ellipse oriented in any direction, or ashape derived from CT images of the heart in the same patient or in anaverage patient. Alternatively, no assumption is made about the shape ofthe heart in processing the impedance data to produce an impedanceimage.

FIG. 7B shows an impedance image 712 of the same cross-section of thechest, derived from impedance measurements gated to the end-diastolephase of the cardiac cycle, when the left ventricle is at its maximumexpansion. The end-diastole phase corresponds to, or correlates to, thepeak of the T-wave in the ECG, which occurs approximately one third of acardiac cycle after the end-systole phase. Optionally, the timing of thenext end-diastole phase is estimated by measuring the ECG signal for atleast a few cardiac cycles, and the impedance measurements are made atthe expected time of the next end-diastole phase.

Optionally, the same method is used to calculate impedance image 712 asis used to calculate impedance image 700, optionally including modelingthe interior of the heart as an ellipse with principal axes 708 and 710.Optionally, in this case, impedance image 700 is used as an initialguess for calculating impedance image 712. Optionally, in calculatingimpedance image 712, the impedance distribution is further constrainedto be the related in some way to impedance image 700, or to be the sameas impedance image 700, except for diameters “a” and “b” of the ellipse,which are allowed to vary to obtain a best fit to the impedance datataken at the end-diastole phase. Alternatively, “a” is constrained to bea particular function of “b,” based for example on empirical datashowing how the shape of the heart changes as it expands. A potentialadvantage of using such constraints is that it may be possible toconverge more quickly on a relatively accurate measure of the volume ofthe interior of the heart, and hence on the stroke volume.

Alternatively, the rest of the chest is not assumed to have the sameimpedance distribution in image 712 as it does in image 700, since thetwo images may be taken at different phases in the breathing cycle.Optionally the change in the impedance distribution of the rest of thechest during the breathing cycle is modeled, and provides constraintsused in calculating impedance image 712. Other methods of dealing withthe effects of the breathing cycle are described below.

Alternatively, impedance data for image 712, at the end-diastole phase,is taken first, and impedance image 700, using data taken at theend-systole phase, is calculated by constraining the impedance image tobe the same except for the heart diameters “a” and “b,” or at leastimage 712 is used as an initial guess for calculating image 700. Apotential advantage of using the end-systole image as an initial guessfor the end-diastole image is that the shape and volume of the interiorof the heart is likely to be more constant at the end-systole phase thanat the end-diastole phase, especially for patients with an irregularheart beat.

Once images 700 and 712 have been found, additional impedancemeasurements are optionally made at the end-systole and/or at theend-diastole phase, using any of the previous images singly or incombination as initial guesses for calculating additional impedanceimages. This done, for example, to improve accuracy by averaging overseveral cardiac cycles and/or several breathing cycles. Optionally,impedance measurements are also made during the expansion and/or duringthe contraction of the heart, to verify that impedance data for images700 and 712 really was taken near the end-systole and end-diastolephases. However, such images, made when the volume of the heart ischanging, may be less accurate than images 700 and 712.

Optionally, measures are taken to insure that the lung volume is notvery different between impedance images 700 and 712, so that it isrealistic to treat the two impedance images as identical except fordiameters “a” and “b” of the heart. For example, the impedancemeasurements for images 700 and 712 are gated to be at a same phase inthe breathing cycle. or the patient holds his breath while the impedancemeasurements are made. Alternatively, several sets of impedancemeasurements are made at different phases in the breathing cycle, andtwo impedance images are chosen that are from nearly the same phase inthe breathing cycle. Alternatively, both impedance images are made usingdata that is averaged over a breathing cycle. Alternatively, theimpedance measurements for image 712 are made at the end-diastole phaseimmediately following the end-systole phase at the measurements forimage 700 are made, close enough in time so that the lung volume hasn'tchanged very much, possibly choosing a time in the breathing cycle whenthe lungs are near their maximum or minimum volume.

The area A of the ellipse used to model the cross-sectional area of theinterior of the heart is given by A=πab/4. The volume V is thenestimated by using the formulaV=8A ²/(3πL)described by H. T. Dodge and F. H. Sheehan, “Quantitative contrastangiography for assessment of ventricular performance in heart disease,”J. Am. Coll. Cardiol. 1, 73 (1983). The estimated stroke volume is thenthe change in the volume V between the end-systole phase, using thevalues of “a” and “b” in image 700, and the end-diastole phase, usingthe values “a” and “b” in image 712.

Alternatively, if a three-dimensional impedance map is reconstructed,for example from impedance data taken from electrodes that are not allin the same plane, then the effective length of the interior of theheart is optionally estimated from the impedance map, rather thaninferred from area A, and the volume V is found by multiplying A by theeffective length.

FIG. 8 shows a bio-impedance system 800, used to obtain impedance imagesof the chest, to evaluate fluid in the lungs, or stroke volume of theheart, for example, in a non-invasive and relatively rapid way. The unitincludes a chest belt 802 with eight electrodes 804. Alternatively adifferent number of electrodes is used, for example 4, 16 or 32, oranother number, even a number that is not a power of 2. Optionally, theelectrodes are disposable, and all the electrodes are replaced by newelectrodes, inserted into connectors in the belt, for each new patient.Optionally, the chest belt comes in different sizes, appropriate forpatients with different chest sizes, and optionally, the length and/ortension of a given size belt is adjustable within limits. Optionally,the electrodes are precisely positioned around the patient's chest, forexample at equal distances around the chest, and/or with one or more ofthe electrodes at a good position for taking ECG data. Alternatively,the electrodes are not precisely positioned, but the positions of thedifferent electrodes optionally are measured and taken into account inreconstructing the impedance images.

An analog amplification unit 806 optionally generates the currents thatare passed through the electrodes to measure the impedance, and controlsthe voltage measurement sequence. As described above for FIG. 1,optionally only some of the electrodes are used to run current through,while other electrodes are used to measure voltage. Optionally, some orall of the electrodes are used sometimes for running current through,and sometimes for measuring voltage, but not at the same time.Alternatively, at least some of the electrodes are used to measurevoltage while they have current running through them.

System 800 also includes a base unit 808, which optionally includes agraphical display and a DSP unit. Optionally, some or all of thefeatures shown in the impedance imaging system in FIG. 4 are also foundin system 800. Optionally, FIG. 8 illustrates a particular packaging ofthe system shown in FIG. 4, or another system, that is easily portableand suitable for use in hospitals. For example, system 800 weighsbetween 2 and 3 kg, or between 3 and 4 kg, or between 4 and 5 kg, orless than 2 kg or more than 5 kg.

System 800, in addition to making impedance measurements, optionallyuses at least one of the electrodes in electrode belt 802 to obtain ECGdata. The ECG data is optionally processed by a controller, such as asmall computer or RISC chip, located in the base unit or in theamplification unit, to determine the phase of the cardiac cycle in realtime, so that the impedance measurements can be synchronized to thephases of the cardiac cycle.

Optionally, the ECG data is from a single lead. Optionally, the ECGsignal from this electrode is displayed on graphical display in realtime, at low resolution, for example 50 pixels per second. Optionally,the operator synchronizes the impedance measurements with the displayedECG signal manually, for example by pressing a start button on the baseunit.

Additionally or alternatively, the system measures the ECG signal, fromone or more electrodes, at greater time resolution, in order toaccurately trigger the impedance measurement at a phase of the cardiaccycle. For example, the ECG signal is measured at intervals of 1 msec,optionally with 12-bit resolution, and a real time algorithm detects thepeak of the R wave, and triggers the impedance measurementautomatically. Optionally, as described above for FIGS. 7A and 7B, thepeak of the T-wave is also detected or estimated, and used to trigger animpedance measurement, for example for measuring the stroke volume ofthe heart.

Optionally, the impedance measurement, once it is triggered manually orautomatically, is performed by injecting current, for example at 20 kHz,into one or more electrodes. The current is collected at one or moreelectrodes, while the voltage (both amplitude and phase relative to thecurrent) is measured at the same electrodes and/or at other electrodesthat are floating. For example, current is passed successively betweeneach of the 28 possible pairs of two electrodes chosen from the 8electrodes in the belt, while voltage is measured by the other 6electrodes which are floating. This set of measurements may provide allof the useful impedance information available from a system with 8electrodes. The voltage at the electrodes that the current passesthrough may be dominated by the impedance of the skin, and beinsensitive to the impedance of the interior of the chest. Passingcurrent through more than two electrodes may not provide additionalinformation, because it is equivalent to a linear combination of passingcurrent through different pairs of electrodes. Voltage measurements areoptionally taken at 200,000 samples per second, at 12 bit resolution,appropriate for making accurate amplitude and phase measurements of a 20kHz signal.

Alternatively, instead of running current through all possible pairs ofelectrodes, and measuring voltage at each of the other 6 electrodes ineach, current is only run through selected pairs of electrodes, andvoltage is only measured between selected pairs of electrodes, which areparticularly useful for measuring the impedance of the interior of thechest. For example, FIG. 9 shows the 8 electrodes of the belt, numbered1 through 8, arranged around a chest 100. Electrodes 1 and 2 are locatedon the chest, in front, 3 and 4 are located on the right side of thebody, electrodes 5 and 6 are located on the back, and electrodes 7 and 8are located on the left side of the body. Tests have shown that, formeasuring fluid in the lungs, it is particularly useful to injectcurrent between electrodes 7 and 3 while measuring voltage betweenelectrodes 8 and 4, and to inject current between electrodes 8 and 4,while measuring voltage between electrodes 7 and 3. For measuring strokevolume of the heart, tests have shown that it is useful to injectcurrent between electrodes 5 and 1 while measuring voltage at electrodes8 and 3, and between electrodes 8 and 2, and to inject current betweenelectrodes 8 and 4, while measuring voltage between electrodes 1 and 3,and between electrodes 1 and 5. Limiting voltage measurements to thesecases produces results for lung fluid and stroke volumes that are almostas good as if all possible measurements were made, while greatlyreducing the time needed to take data, and the complexity of the dataanalysis. Alternatively, different combinations of electrodes are used.Alternatively, a larger set of measurements is made, but still not usingall possible combinations of electrodes, for example current is injectedthrough each of three different pairs of electrodes, and two voltagemeasurements are made for each pair of current-carrying electrodes.Optionally, only the electrodes that are needed for a particular set ofmeasurements are actually present. For example, if only fluid in thelungs is to be measured, using the set of measurements described abovefor this purpose, then optionally the belt only has four electrodes,corresponding to electrodes 3, 4, 7, and 8 in FIG. 9. Alternatively,instead of spacing these four electrodes as shown in FIG. 9, they arespaced evenly around the chest.

Each measurement preferably lasts for at least several cycles of the ACcurrent, with adequate dead time between measurements, so that accurateamplitude and phase measurements of the voltage can be made. The totaltime of all the measurements is optionally short compared to a cardiaccycle or a breathing cycle, so that all of the measurements are made atthe same phase of the cardiac and breathing cycles. Alternatively oradditionally, the measurements are made over more than one cardiac orbreathing cycle, but are optionally gated to the cardiac and/orbreathing cycles, so that there is enough time to make all themeasurements, and/or to improve accuracy by averaging over more than onemeasurement. Alternatively or additionally, the measurements are madeover a range of phases of the cardiac cycle and/or the breathing cycle,but are binned according to the phase of the cardiac cycle and/or thebreathing cycle. Alternatively or additionally, measurements are takingover a limited range of phases of the cardiac cycle and/or the breathingcycle, but in a part of the cycle where the heart (in the case of thecardiac cycle) or the lungs (in the case of the breathing cycle) arenear a maximum or minimum in volume. Optionally, averaging or other dataprocessing is done in real time while the data is taken.

Optionally, in addition to using the ECG data to trigger the impedancemeasurements, information obtained from the ECG data, for example the RRinterval, the QT interval, and/or the QTc interval, is calculated anddisplayed on the graphical display, and/or included in a data file ofthe impedance data. This information from the ECG may be useful, forexample, in evaluating the health of the patient, and in interpretingthe medical significance of the lung fluid measurements and/or strokevolume measurement obtained from the impedance data.

Optionally, the impedance data is stored, and the impedance images arecalculated later. Alternatively, system 800 includes a computer ofadequate power to calculate impedance images in real time.Alternatively, a computer or a controller in system 800 does not makefull calculations of the impedance image in real time, but analyzes thedata sufficiently to verify that the data is reasonable, and is notaffected, for example, by inadequate contact between one of theelectrodes and the skin, or by a malfunction in the amplification unit,and warns the operator in real time if the data is not good.

Optionally, system 800 has the ability to make a wireless transfer ofimpedance data, and/or results calculated from the impedance data suchas impedance images, for example for archiving purposes. The wirelessdata transfer capability optionally also allows system 800 to receivedata from other devices, for example data about the patient such asweight, that may be used in calculating the impedance images or inevaluating them. The wireless data transfer is done, for example, usingan infra-red transceiver or a wireless modem device. In addition to orinstead of wireless data transfer capability, system 800 also optionallyhas data ports such as an R232 interface, for data input and/or output.Optionally, system 800 prints out certain information obtained from theimpedance measurements or ECG, either through a data port, or using aprinter 810 incorporated in the base unit for example. Optionally, theprinter is an inexpensive and/or light weight and/or low power printer,such as a thermal printer.

System 800 is optionally battery operated, which has the potentialadvantages of low electrical noise and improved safety. Optionally thebattery is rechargeable, and optionally system 800 can run for at least5 hours before recharging or replacing the battery. Optionally, thesystem has an “idle” mode which reduces power consumption by turning offpower consuming elements that are not being used. A low batteryindicator is optionally included.

The graphical display optionally has touch screen capability. Forexample it is a touch screen LCD with a black and white graphicaldisplay, with 320 by 240 pixels. Alternatively a different kind ofgraphical display, or a color display, or a display with a differentnumber or arrangement of pixels is used. Optionally there are nooperating buttons for the system except the touch screen.

FIG. 10 is a flow chart showing an exemplary procedure for using system800. In 902, the operator makes sure that the patient does not have anypace-maker or defibrillator implants, since even the low currents usedfor the impedance measurements could dangerously affect such implants.If the patient does not have such implants, then the patient's chest ismeasured, in 904, to determine which size belt to use. In 906, thedisposable electrodes are inserted into the proper connectors in thechest belt. In 907, the patient's chest is cleaned, and in 908 the chestbelt with the electrodes is applied to the chest. In 910, the ECG signalis monitored, to verify that the electrodes are properly placed, and ifnot, their positions are adjusted, and/or the tension of the belt isadjusted, in 912, and the ECG signal is monitored again, until theelectrodes are properly placed.

In 914, the sequence for taking impedance data is initiated, for exampleby pressing a start key on the touch screen. The sequence comprises, forexample, the six voltage measurements made while sending current througheach of the 28 pairs of electrodes, or a useful subset of thosemeasurements, as described above. Optionally, the whole sequence isautomated, including triggering by the ECG signal, without the need forany intervention by the operator. Optionally, at least a preliminaryanalysis is made of the data in 916, and if there is an indication thatthe data is not good, due for example to a malfunction of the system orimproper placement of the electrodes, the system indicates this andoptionally diagnoses the problem, for example by a message on thegraphical display and/or by an audible tone such as a beep. The operatorthen optionally attempts to correct the problem in 918, and initiatesthe data taking sequence again.

When the sequence is successfully completed, as indicated for example bya message on the graphical display, and/or an audible signal, then atleast a portion of the results are optionally printed out, in 920, forexample by pressing a print key on the touch screen. Alternatively oradditionally, some or all of the results are displayed on the graphicaldisplay. If some results require post-processing on a computer notincluded in system 800, then the data needed for the post-processing istransferred to the computer through a data port in 922, eitherautomatically at the end of the data taking sequence, or when requestedby the operator, for example by pressing a button on the touch screen.Alternatively, this data transfer is done before printing out theresults, or at the same time. The chest belt is removed in 924, thedisposable electrodes are removed from the belt, and the patient's chestis cleaned.

The word “data analyzer” as used herein means any equipment used toanalyze data, even if it is not a single unit. For example, when a dataanalyzer is described as analyzing electrocardiograph data andreconstructing an impedance image, this does not necessarily mean that asingle piece of equipment does both the analyzing and thereconstructing. The word “data analyzer” can include one or moreordinary computers running software, one or more pieces of speciallydesigned hardware, or both. The words “comprise”, “include” and theirconjugates as used herein mean “include but are not necessarily limitedto”. While the invention has been described with reference to certainexemplary embodiments, various modifications will be readily apparent toand may be readily accomplished by persons skilled in the art withoutdeparting from the spirit and scope of the above teachings.

What is claimed is:
 1. A method for measuring fluid in a subject'slungs, comprising: a) placing at least four electrodes around asubject's chest; b) obtaining one or more sets of impedance data, eachset consisting of a voltage measured between a first pair of theelectrodes while passing current through the subject's chest through asecond pair of the electrodes, both the electrodes in the second pairnot belonging to the first pair; and c) estimating a degree of fluid inthe subject's lungs, using one or more of the sets of impedance dataobtained from a total of eight or fewer of the electrodes.
 2. A methodaccording to claim 1, wherein, for at least first one of the sets ofimpedance data, one of the pairs of electrodes is placed respectively onthe front right side and back left side of the subject's chest, and theother one of the pairs of electrodes is placed respectively on the frontleft side and back right side of the subject's chest.
 3. A methodaccording to claim 2, wherein at least a second set of impedance data isobtained, for which the voltage is measured between the pair ofelectrodes through which the current is passed for the first set, whilethe current is passed through the pair of electrodes between which thevoltage is measured for the first set.
 4. A method according to claim 2,wherein the pairs of electrodes for the first one of the sets ofimpedance data are placed closer to the sides of the subject's body thanto a midplane of the subject's body.
 5. A method according to claim 1,wherein estimating the degree of fluid in the subject's lungs comprisessolving an inverse problem to find a specific impedance of the subject'slungs, according to an image or model of the chest.
 6. A methodaccording to claim 1, also comprising obtaining electrocardiograph dataof the subject, wherein estimating the degree of fluid in the subject'slungs comprises: reconstructing a plurality of preliminary impedanceimages of the chest from the electrical data; using theelectrocardiograph data to determine a phase of breathing cycle of thesubject, a cardiac cycle of the subject, or both, for each of thepreliminary images; sorting the preliminary images into a plurality ofbins according to the cardiac cycle, the breathing cycle, or both; andusing the sorted preliminary impedance images to obtain a correctedimpedance image of the subject's chest.
 7. A method according to claim5, wherein the image or model of the chest is used to make an initialguess for an impedance distribution of the chest, for solving theinverse problem.
 8. A method according to claim 5, wherein the image ormodel of the chest is used to make a grid, used for solving the inverseproblem, conform to a surface of the lungs, a surface of the heart, orboth.
 9. A method according to claim 5, wherein solving the inverseproblem comprises using a finite volume method and calculating aJacobian analytically.
 10. A method according to claim 1, whereinplacing the electrodes around the chest comprises placing all theelectrodes substantially at a same level around the chest.
 11. A methodaccording to claim 1, wherein the current is between 1 and 5 milliamps.12. A method according to claim 1, wherein the current is an alternatingcurrent at a frequency between 10 kHz and several hundred kHz.
 13. Amethod according to claim 1, wherein placing the electrodes around thechest comprises placing around the chest a belt that the electrodes aremounted on, the belt being placed tightly enough to hold the electrodesagainst the chest.
 14. A method according to claim 13, also comprisingusing at least some of the electrodes mounted on the belt to obtainelectrocardiograph data of the subject, while obtaining the one or moresets of impedance data.
 15. A method according to claim 1, wherein thetotal of eight or fewer of the electrodes consists of exactly fourelectrodes.
 16. A method according to claim 1, wherein the total ofeight or fewer of the electrodes consists of exactly five electrodes.17. A method according to claim 1, wherein the total of eight or fewerof the electrodes consists of exactly six electrodes.
 18. A methodaccording to claim 1, wherein the total of eight or fewer of theelectrodes consists of exactly seven electrodes.
 19. A method accordingto claim 1, wherein the total of eight or fewer of the electrodesconsists of exactly eight electrodes.
 20. A method according to claim 1,wherein each of the one or more sets of impedance data is obtained whilepassing current through only one pair of the electrodes.
 21. A systemfor measuring fluid in a subject's lungs, comprising: a) at least fourelectrodes adapted for placing around a human subject's chest; b) apower supply capable of passing a known current through the subject'schest through a first pair of the electrodes, when they are placed onthe subject's chest; c) a voltage measuring instrument capable ofmeasuring a voltage between a second pair of the electrodes, distinctfrom the first pair, when the power supply is passing the known currentthrough the subject's chest through the first pair of electrodes, andwhen the first and second pair of electrodes are placed on the subject'schest; and d) a controller which estimates a degree of fluid in thesubject's lungs, using one or more voltages measured between pairs ofthe electrodes for known currents passed through other pairs of theelectrodes, for electrodes placed at no more than eight locations aroundthe chest.
 22. A system according to claim 21, wherein the controller isconfigured to estimate the degree of fluid in the subject's lungs bysolving an inverse problem to find a specific impedance of the subject'slungs, according to an image or model of the chest.
 23. A systemaccording to claim 21, comprising a belt that the electrodes are mountedto, which holds the electrodes in place against the chest when it isplaced around the chest.
 24. A system according to claim 21, comprisingan electrocardiograph, wherein the controller is configured toreconstruct a preliminary impedance image of the chest using thevoltages measured for the known currents, to reconstruct a correctedimpedance image from the preliminary impedance image using data of thesubject from the electrocardiograph to correct for the subject's cardiaccycle, breathing cycle, or both, and to estimate the degree of fluid inthe subject's lungs from the corrected impedance image.
 25. A systemaccording to claim 21, wherein the power supply comprises an AC powersupply capable of passing at least 1 milliamp of current at a frequencyof at least 10 kHz.